The administration of aerosolised medicines is one of the main pillars of therapy for a number of pulmonary diseases such as asthma, chronic obstructive pulmonary disease (COPD), infant respiratory distress syndrome (IRDS), pulmonary arterial hypertension (PAH) or cystic fibrosis (CF). The particular advantage of inhalation therapy is that, in principle, the aerosolised medicine is directly administered to the affected organ, i.e. the respiratory system, rather than to the systemic blood circulation from where the drug substance is distributed to the lungs, but also to other organs and tissues where the compound is not desired and can cause side effects.
However, specific delivery of aerosols by means of the inhalation route of administration is not straight-forward. The degree to which an aerosol reaches its target destination within the respiratory system depends on numerous factors, including the aerosol parameters, such as the aerodynamic diameter of the aerosolised particles or droplets, which result from the pharmaceutical composition and the inhalation devices that it used to convert the composition into an inhalable aerosol; but also on patient-related factors, such as the volume of inhalation and, in particular, the inspiratory flow rate and/or the inspiratory pressure. For example, many patients inhale at too high inspiratory flow rates and/or underpressures, assuming that “sucking in” their medication would deliver higher fractions of it into the deep lungs, and thus be most beneficial. Tiddens et al. (Journal of Aerosol Medicine; Vol. 19, Nr. 4, 2006; Pp. 456-465) describes that even patients suffering from cystic fibrosis achieve a mean peak inspiratory flow rate (PFI) of 52 L/min (range 26-70) for all patients (47 L/min (26-62) in children) at the highest resistance tested for a dry powder inhaler; minimal flow rates of 30, 45, and 60 L/min at that highest resistance were obtained in 99%, 80%, and 22% of all patients. While such high flow rates may be necessary in dry powder inhalers to disperse the powdered formulation, they are not desirable for other inhalation devices which emit a pre-dispersed aerosol and they increase the likelihood that aerosol particles or droplets impact in the throat of a patient rather than the lungs. Throat deposition not only means that the respective fraction of the therapeutic compound is lost, but also an increased risk of systemic or local side effects. In addition, when inhaling too fast through a device providing flow resistance, as is e.g. already common for dry powder inhalers, the underpressure increases. High underpressure in a patient's lung will induce a progressive constriction of small alveoli and bronchi, with the consequence that less drug can be deposited there.
Therefore, attempts are being made to train patients to perform a proper breathing manoeuvre in order to achieve appropriate deposition of the aerosol formulation. However, such training is often perceived as inconvenient by the patient and/or the health professional. Moreover, the capability of patients of complying with breathing instructions is very diverse. In particular children, elderly patients, patients with motoric difficulties or mental limitations are often not able to perform a breathing manoeuvre correctly as instructed. Thus the success of training alone is limited.
Furthermore, these issues become even more important when the number of breaths required to administer a desired dose increases. While, for example, many dry powder inhalers, pressurised metered dose inhalers and/or soft mist inhalers deliver the required dose within just one or two inhalation manoeuvres upon actuation, drug administration with most nebulisers, such as ultrasonic nebulisers, jet nebulisers and/or vibrating mesh nebulisers typically involves a larger number of inhalation manoeuvres and thus longer administration times. The risk of deviating from a trained breathing manoeuvre increases with administration time; e.g., due to distraction or getting less concentrated. Furthermore, an underpressure which may be well tolerable for one or two breaths only, e.g., inhaling at −40 mbar with a dry powder inhaler, can become rather inconvenient and exhausting when more inhalation manoeuvres are required.
To compensate for patients' difficulties and limitations, improved inhalation devices have been developed that take the capabilities of individual patients into account and provide aerosols with optimised flow and volume. For example, the AKITA® JET inhalation system, which is based on a conventional jet nebuliser in combination with a control unit that actively regulates inspiration flow as well as inhalation volume, substantially improves the deposition of an aerosolised drug in the target region compared to purely patient-controlled inhalation. Using such system also leads to a more reproducible lung deposition and thus to a more predictable therapeutic effect.
One of the limitations of such systems which control flow and volume is that they are not readily portable. It is difficult to implement the control features and functionalities within a small handheld device, which is the type of inhaler that is preferred by some patients due to its portability. Some attempts to ensure, or at least facilitate, appropriate aerosol inhalation manoeuvres, even with hand-held devices, have been described in the prior art.
EP 2 283 887 B1 discloses a miniaturised device for variable flow rate limitation at low differential pressure (or, in case of an inhalation, underpressure, or negative pressure), in particular for the limitation of the inhalation flow during the inhalation of therapeutic aerosols. While EP 2 283 887 B1 is silent as to any specific type of inhalation device such as powder inhalers, pressurised metered dose inhalers or nebulisers in which it is to be used, the device is small enough to be accommodated e.g., in a hand-held inhalation device. Depending on the geometric dimensions of the device and/or the flexibility of the membrane employed in it, among other parameters, said device for flow rate limitation (or flow rate restrictor, or flow restrictor) provides a variable flow restriction; i.e. the flow rate does not increase linearly with increasing underpressure; in other words, at a low underpressure, there is relatively less flow restriction than at a high underpressure. Depending on its configuration, such variable flow restrictor may even provide for a maximum flow rate; i.e. even if the patient tries to inhale faster and thus increases the underpressure, the flow rate does not increase much further beyond a maximum flow rate.
However, it is difficult and undesirable to rely solely on flow restriction in order to ensure a desirable low inspiratory flow rate. For example, to prevent a patient from inhaling at a flow rate of more than about 12 to 18 L/min, a flow restrictor with substantial flow resistance would have to be used, which especially patients suffering from obstructive airway diseases such as asthma or COPD may consider rather uncomfortable. Hence, it can be necessary to choose a lower flow resistance for more severely affected patients at the cost of not being able to prevent patients from using higher flow rates, e.g. up to about 25−30 L/min instead of a desired flow rate in the range of e.g., 15 L/min. Moreover, patients with less compromised pulmonary functions may intuitively react to a high flow resistance by further increasing the underpressure with which they generate inspiratory flow, causing the pressure in the device, which should preferably be not lower than about −20 mbar, to decrease to values of about −30 mbar or less. Inherently, the pressure in the patient's lungs would be even lower. As mentioned above, such pronounced underpressure is undesirable because it induces a progressive constriction of small alveoli and bronchi, with the consequence that less drug can be deposited there. Since the inhalation devices equipped with a variable flow restrictor according to EP 2 283 887 B1 do not provide any control of the underpressure, this approach does not ensure proper, reproducible inhalation manoeuvres. In addition, EP 2 283 887 B1 also does not disclose any control of other inspiratory parameters such as the inhaled volume or inspiration time.
WO 2011/083377 A1 describes a feedback and compliance device which may be coupled to a pressurised metered dose inhaler (pMDI), dry powder inhaler (DPI) or an aqueous liquid dispensing system. Nebulisers such as ultrasonic nebulisers, jet nebulisers or vibrating-mesh nebulisers are not disclosed. The device comprises sensors to sense parameters relating to the use of the apparatus, and a processing unit which is programmed to cause one or more feedback device(s) to provide information to a patient based on said sensors' output. The parameters relating to the use of the apparatus for which feedback is provided are e.g.,                i) correct insertion of the medication storage into the apparatus,        ii) proper shaking of the drug delivery apparatus,        iii) start of inhalation in an appropriate time after shaking,        iv) inhalation for a proper time period, and/or        v) breath hold after inhalation.        
According to the flowcharts provided in the document, the preset feedback typically follows a simple, binary yes-or-no-decision, and is specifically adapted for the correct use of pressurised metered dose inhalers.
WO 2011/083377 A1 emphasises that the feedback device does not modify or interfere with the flow introduced by actuation of the medication storage. There is no provision for controlling inspiratory parameters such as inspiratory flow rate and/or underpressure.
US 2011/0226242 A1 describes a respiratory drug delivery apparatus that includes a housing for holding a source of medication, a valved holding chamber/spacer from which the patient inhales and an audible feedback device coupled thereto. The audible feedback device is a sound generator adapted to generate audible instructions in response to sensor signals caused either by manual actuation (e.g., pressing a button, removing a cap, inserting a source of medication into the housing) or in response to an event relating to the operation of the apparatus (e.g., actuation of a pressurized metered dose inhaler (pMDI), opening of a valve in the spacer, etc.). Said audible instructions may comprise pre-recorded instructions relating to shaking and actuating the pMDI, proper inhalation and holding one's breath, which either correct or prevent an incorrect use or reinforce the correct one. For example, the device may be configured to provide audible instructions for taking and counting a certain number of breaths or for inhaling for a particular period of time. Alternatively, the audible feedback device may be a noisemaker, such as a whistle or sound reed integrally formed in the holding chamber, which may start to emit a sound when the patient needs to be instructed with respect to the next action to take, or alerted to an event or a problem.
US 2011/0226242 A1 is silent about important parameters relating to the breathing manoeuvre which have substantial impact on the degree and site of drug deposition, in particular desirable ranges for inspiratory flow rates and inhaled volumes. While the document also mentions the possibility of using audible feedback to a patient to encourage slow inhalation, it is highly unlikely that such general instruction will enable a patient to actually achieve inhalation at a particular target flow rate and/or underpressure.
US 2005/087189 A1 discloses a device for the delivery of drug-laden or drug-free air during multiple inhalations. The drug delivery device comprises sensors for monitoring the breathing pattern of a patient; a processor with an internal clock to analyse said breathing pattern in order to control the onset and duration of the drug-laden and calculate the cumulative, administered dose; and a feedback indicator. The feedback indicator provides information to the patient as to whether the monitored breathing pattern is effective or suitable for inhaling drug-laden air. If the breathing pattern is too weak or too unsteady, the drug-laden pulse will not be delivered or stop early. For this purpose, the device comprises sensors which detect the onset and end of an inhalation. The device may further comprise a sensor for detecting the introduction of drug into a holding chamber.
However, US 2005/087189 A1 does not disclose any intention or means for controlling certain important parameters relating to the breathing manoeuvre, in particular a desirable inspiratory flow rate, inhaled volume, and/or underpressure.
U.S. Pat. No. 5,906,202 A describes a hand-held, self-contained and readily portable respiratory drug delivery device which is capable of measuring and recording the patient's total respiratory tract capacity, the inspiratory flow rate and inspiratory volume, e.g. by a microprocessor in combination with a read/write memory means and a flow measurement transducer. These spirometric parameters are measured in a monitoring event before the actual dosing event (also called drug delivery event) in order to calculate the time point when, during a patient's inhalation phase, to actuate the release of an aerosol bolus into the inspirational air flow, and the volume of said aerosol bolus. The actuation occurs automatically and leads to the spring-driven ejection of one or more doses of liquid drug formulation from a blister strip through a porous membrane with defined pore sizes. After the release the aerosol bolus and its inhalation by the patient, the flow can be shut off completely or partially by means of a ball valve, needle valve, gate valve or pinch valve. Alternatively, the device may provide a signal (light or sound) to the patient requesting him/her to stop inhalation.
The device may further comprise visual feedback components such as light diodes which display to the patient in a countdown fashion the remaining seconds during which breath must be held. It may also prompt the patient to hold his breath until notified by a visual signal (e.g., flashing light) or an audio signal. The device may be equipped with a flow rate sensor, and visual signals may indicate to the patient whether or not he inhales at the preferred rate.
However, it is difficult for patients to achieve a desirable low flow rate of e.g. 12 to 18 L/min on the basis of this type of feedback alone. Rather, most patients would tend to produce a substantially fluctuating inspiratory flow rate, leading to variable drug deposition and unpredictable therapeutic effect. The feedback system requires the patient to be fully concentrated on the feedback signals, which may be difficult in particular for children and elderly patients, but also for most other patients under treatment regimen which require long inhalation times.
Thus, there is an ongoing need for inhalation devices that overcome one or more of the limitations of the presently known devices. In particular, there is a need for improved inhalation devices which make it easier for patients from very different patient populations and with rather different levels of pulmonary function to perform breathing manoeuvres optimised for a specific therapeutic application, to achieve and maintain a desired inspiratory flow rate and/or to avoid applying too much underpressure when inhaling a therapeutic aerosol.
These needs are addressed by the present invention whose object is to provide such improved devices. Other needs and objects of the invention will become clear on the basis of the description and the patent claims.